This invention pertains generally to the field of microscopy and spectroscopy, and particularly to laser scanning fluorescence microscopy.
Scanning optical microscopes, such as laser scanning confocal microscopes, are of increasing importance in microscopy, particularly for imaging of dynamic biological structures such as living cells. In a scanning microscope, the light beam from the source, usually a laser, is focussed to a point within the specimen by the microscope objective and the specimen and beam are moved relative to one another in a raster fashion, either by moving the stage on which the specimen is mounted or, more commonly, by deflecting the light beam so that it scans across a stationary specimen. The light from the specimen is collected by the objective and passed back through the microscope to a detector, such as a photomultiplier tube. In addition to detection of light reflected from a specimen (or transmitted through the specimen), scanning microscopes can also be constructed to detect fluorescence induced by the illuminating light beam. Typically, the fluorophores in the specimen absorb the illumination light, which is at a chosen wavelength (usually shorter wavelength visible light), and fluorescently emit photons at a longer wavelength which are received by the objective of the microscope and passed back through the scanning optics to a dichroic mirror which separates the fluorescent light from light at the illuminating light wavelengths and directs the fluorescent light to a separate photodetector. In this manner, particular structures within the specimen, such as parts of cells, can be labeled with fluorescent markers and distinctively imaged by the scanning microscope.
Most fluorophores can also absorb two (or more) photons of longer wavelengths simultaneously when sufficiently intense illumination light is applied thereto and will emit a fluorescent photon at a shorter wavelength than the incident light. This phenomenon is exploited in multi-photon laser scanning microscopes in which an incident beam of relatively long wavelength light in short pulses from a laser source is narrowly focussed onto a specimen so that the light reaches an intensity at the focal point sufficient to excite detectable two (or more) photon fluorescence. The emitted fluorescent photons collected by the objective lens of the microscope are passed back through the optical system, either through the scanning optics to a dichroic mirror which reflects light at longer wavelengths while passing the shorter wavelength fluorescent light to a separate detector, or by bypassing the scanning system, and directing the light from the microscope objective lens to a dichroic mirror which passes the shorter wavelength fluorescent light directly to a detector while reflecting the longer wavelength excitation light. See, Winfried Denk, et al., xe2x80x9cTwo Photon Laser Scanning Fluorescence Microscopy,xe2x80x9d Science, Vol. 248, Apr. 6, 1990, pp. 73-76; Winfried Denk, et al., xe2x80x9cTwo-Photon Molecular Excitation in Laser-Scanning Microscopy,xe2x80x9d Chapter 28, Handbook of Biological Confocal Microscopy, Plenum Press, New York, 1995, pp. 445-458; and U.S. Pat. No. 5,034,613 entitled Two-Photon Laser Microscopy. By focussing the incident light from the objective lens to a relatively narrow spot or beam waist such that the intensity of the incident light is sufficient to excite multi-photon excitation only at the waist within the specimen, multi-photon fluorescence excitation will occur generally only in the focal plane. The shorter wavelength fluorescent light emitted by the specimen can then be passed back, either through the scanning system to de-scan the light or directly, without de-scanning, to a fluorescent light detector to obtain an image corresponding to the focal plane. Therefore, the excitation light alone produces the desired depth resolution (i.e., an optically sectioned fluorescence image), so that there is no need for the use of a confocal aperture.
Fluorescent signal photons can be characterized by their wavelength, the lifetime of the excited state giving rise to the photon, and polarization. These parameters can be used to identify a fluorophore or to provide information on the microenvironment of the fluorophore. Fluorescence microscopy of living specimens generally yields very weak signals. Therefore, any multi-dimensional spectral imaging system must be very photon efficient to be practical for in vivo imaging. In addition, such systems must have the lowest possible values of intrinsic (i.e., system generated) noise. Living cell studies, such as four dimensional imaging or ion imaging, generally require faster imaging speeds than are currently available from commercial multi-photon laser scanning microscopes (MPLSM). Several fast scanning MPLSM systems are currently in use in research laboratories, but these instruments either do not preserve the deep section contrast advantages of multi-photon over confocal microscopy or do not allow use of electronic magnification of the scanned area.
Most biological tissue is autofluorescent. Molecules such as NAD(P)(H), elastin, and chlorophyll act as endogenous fluorophores. These endogenous fluorophores can often be identified by their characteristic spectra. A spectral imaging system would thus be of considerable use in identifying endogenous fluorophores and specifying spectral windows that would either maximally accept or reject these signals, depending on the application.
The use of engineered fluorescent probes as physiological indicators has become a well established technique. Some probes indicate the presence of a bound ligand by changes in fluorescence intensity (e.g., calcium Green 1) while others use spectral shifts (e.g., Indo 1). The latter are favored because ratio imaging at two different wavelengths may be used to provide measurements that are independent of the concentration of the indicator molecule, requiring quantitative measurements. However, each of these probes now requires the use of a custom two-channel filter set.
Fluorescence resonant energy transfer (FRET) is a powerful technique for measuring intermolecular distances in vivo. This technique also now requires custom filter sets that are matched to the donor and the acceptor molecule""s emission spectra. Ratiometric measurements are used to measure the extent of resonance transfer. This technique is proving to be valuable for the in vivo visualization of the docking of a receptor with its ligand, and is the basis of operation of a GFP based calcium indicator.
Fluorescence in situ hybridization is another significant area where multiple fluorophores and ratiometric techniques are used. Often, the main requirement in this application is to spectrally resolve as many separate fluorescent probes as possible.
A major consideration in the detection of fluorescence from scanning microscopes is the ability to collect the desired signal in the presence of significant noise (detection noise, system noise, fluorescent background, etc.). Background fluorescence from endogenous fluorophores or from another interfering exogenous fluorophore can severely reduce detection, or interpretation, of the image signal. In samples labeled with multiple fluorophores, the signal from one fluorophore is often much stronger than another and can spill over to an adjacent channel. In such instances, it is often necessary to move the spectral detection windows as far apart as possible to aid discrimination between the two fluorophores being studied rather than choosing spectral windows to give the maximum signal in each channel. The use of multiple fluorescent labels has been commonplace in the study of fixed specimens, and is now being established for use in in vivo studies. There are now many fluorophores that are available, each with its own unique spectral characteristics. The large number of available fluorophores has carried with it the problem that many different filter sets are now required for double or triple labeled samples. Filter sets use expensive interference filters and dichroic mirrors, and are often difficult to interchange. An ideal filter would have to be continuously adjustable so that for any particular combination of fluorophores used, an optimal set of band-pass assignments could be selected for each detection channel to minimize signal bleed-through and maximize the signal-to-noise ratio.
There are several major advantages that a spectral imaging system would have over a conventional, filter-based three-channel detector. These include dynamic identification of auto-fluorescence and optimization of windows for rejection or imaging as required; dynamic optimization of spectral windows for multiple labels; dynamic background subtraction of reference spectra before the image is even displayed; identification of spectral shifts of fluorophores in different environments; full signal optimization for any ratiometric indicator; fluorophore separation after data collection if a full spectral image is taken can be carried out; and evaluation of standard histological procedures for MPLSM analysis is permitted for identification of tissue-specific spectral shifts of staining.
There are two types of spectral imaging systems currently available commercially. One type uses a Fourier transform spectrometer attached to a standard fluorescence microscope. These systems (an example of which is marketed by Applied Spectral Imaging) split the imaging path in a Sagnac interferometer. Successive interferograms are obtained as the optical path length of one arm of the spectrometer is scanned. Spectral information is recovered by reverse Fourier transformation of the captured stack of interferograms. This method can give high resolution spectral information at each pixel of an image, but has the disadvantage that many interferograms have to be collected, each requiring a separate image. Thus, many frames must be captured with a consequent high level of fluorophore exposure to excitation. In addition, the technique is computationally intensive and requires several minutes to collect data and to compute a spectrogram. Because spectral image reconstruction is so slow, the technique is only suitable for fixed specimens. A further limitation for 3D MPLSM systems is that well collimated light is needed for the interferometer. Scattered emission light that could normally be used for MPLSM imaging cannot be used with this type of imaging spectrometer. A second type of spectral imaging system, available commercially from Leica (the Leica SP) is a confocal microscope that features adjustable spectral windows for signal detection. The system has four detection channels that can be assigned to four arbitrary, non-overlapping regions of the spectrum, and uses a simple prism spectrometer and a system of adjustable, reflecting slits to select parts of the spectrum to four photomultiplier tubes. Light impinging on the sides of one slit is deflected into an adjacent photomultiplier. This scheme is efficient but is limited to four channels, and is thus not a spectral imaging system as such. More channels would be required to visualize the shape of a spectrum so as to identify fluorophores (such as sources of background fluorescence) on the basis of their characteristic emission spectra. The prism spectrometer utilized in the Leica SP capitalizes on the highly collimated signal light path in a conventional confocal microscope. However, there are advantages in positioning the detectors in a MPLSM imaging system so that the maximum amount of scattered (non-collimated) light is also collected for signal detection. See Wokosin, et al., xe2x80x9cDetection Sensitivity Enhancements for Fluorescence Imaging With Multi-Photon Excitation Spectroscopy,xe2x80x9d Proc. 20th Int""l. Conf. IEEE Engr. In Medicine and Biology Soc., Vol. 20, No. 4, 1998. This scattered light collection technique cannot be utilized with a de-scanned emission prism spectrometer using high F-number optics, and thus a key advantage of MPLSM, enhanced emission collection and deeper imaging, would not be readily obtained with this type of spectrometer.
Time-resolved fluorescence spectroscopy is a well-established technique for studying the emission dynamics of fluorescent molecules, i.e., the distribution of times between the electronic excitation of a fluorophore and the radiative decay of the electron from the excited stated producing an emitted photon. The temporal extent of this distribution is referred to as the fluorescence lifetime of the molecule. Lifetime measurements can yield information on the molecular microenvironment of a fluorescent molecule. Factors such as ionic strength, hydrophobicity, oxygen concentration, binding to macromolecules and the proximity of molecules that can deplete the excited state by resonance energy transfer can all modify the lifetime of a fluorophore. Measurements of lifetimes can therefore be used as indicators of these parameters. Furthermore, these measurements are generally absolute, being independent of the concentration of the fluorophore. This can have considerable advantages. For example, the intracellular concentrations of a variety of ions can be measured in vivo by fluorescence lifetime techniques. Many popular, visible wavelength calcium indicators, such as Calcium Green, give changes of fluorescence intensity upon binding calcium. The intensity-based calibration of these indicators is difficult and prone to errors. However, many dyes, such as calcium crimson, exhibit useful lifetime changes on calcium binding and therefore can be used with lifetime measurements. This gives the considerable advantage that absolute measurements of concentration can be made with no elaborate calibration procedures required. Alternatively, lifetime measurements may be used to calibrate the intensity signals from these indicators when maximum sensitivity is required.
An exciting new development has been the technique of fluorescence lifetime imaging microscopy (FLIM). In this technique, lifetimes are measured at each pixel and displayed as contrast. Lifetime imaging systems have been demonstrated using wide-field, confocal and multi-photon imaging modes. FLIM combines the advantages of lifetime spectroscopy with fluorescence microscopy by revealing the spatial distribution of a fluorescent molecule together with information about its microenvironment. In this way, an extra dimension of information is obtained. This extra dimension can be used to discriminate among multiple labels on the basis of lifetime as well as spectra. This would allow more labels to be discriminated simultaneously than by spectra alone in applications where many labels are required, such as FISH, for example. There are also promising applications of lifetime imaging in the medical sciences. For example, tumors have been detected in mice sensitized with a hematoporphorin derivative by lifetime imaging.
Multi-photon lifetime imaging of live specimens is particularly interesting. In these applications, lifetime imaging in conjunction with spectral imaging, should greatly facilitate studies using ion indicator probes and FRET studies of intermolecular distances. Lifetime measurements are a sensitive indicator of FRET and, in combination with spectral measurements, should provide a more sensitive indication of, for example, calcium levels when used with FRET based calcium indicators such as cameleon.
Lifetime measurements are a means of providing another dimension of information from fluorescent probes used in vivo. In most applications where probes are viewed in four-dimensions in vivo, there is a benefit from more or better information. Fluorescent lifetimes can be measured either in the frequency domain or in the temporal domain. Three general strategies have been used to measure fluorescence lifetimes.
A first technique is frequency domain imaging in which a high-frequency, modulated light source is used for fluorophore excitation. By the use of a gain-modulated detector, the phase shift and amplitude demodulation of the fluorescence signal is determined. From these data, the fluorescent lifetime of the probe can be calculated. This scheme is robust and has been extensively used. However, it suffers from several drawbacks: the detector is only working at 50% of its maximum efficiency because it is gain modulated, several data sets taken at different excitation modulation frequencies have to be taken in order to separate two or more lifetime components and, finally, this scheme does not work well with photon counting techniques.
A second technique is time-domain lifetime imaging with gated detector in which a gated micro-channel plate image intensifier is used in conjunction with a CCD imaging camera. Spectral information is obtained by gating the image intensifier on for a narrow time-window at progressively later intervals after the excitation pulse in a succession of data frame captures. This scheme is probably the simplest way of implementing a lifetime imaging system. However, it suffers from two major drawbacks. The method has very poor photon utilization as only one temporal interval is detected at a time. If there are 32 intervals, for example, {fraction (31/32)} of the signals is not utilized and 32 separate frames have to be captured. The second reason this scheme is not appropriate for a multi-photon imaging application is that an imaging (i.e., area) detector is used. This means that the deep sectioning advantage of multi-photon imaging is not realized because scattered fluorescence emission photons will give rise to background noise rather than contributing to the signal, as can be done with a point-scanning multi-photon system.
A third technique is time-domain lifetime imaging with photon counting. At low-light levels, photon-counting detectors have considerable advantages in that they can virtually eliminate noise contributions from electronic amplifiers or electron multiplier noise in a photomultiplier. Also, photon-counting systems provide quantized pulses for every detected photon, allowing the lifetimes to be measured directly using electronic circuitry. Because of the very high speeds necessary to obtain sub-nanosecond temporal resolution, time-to-voltage converters are usually used to measure the interval between the fluorophore excitation pulse and the time of detection of the emitted fluorescent photon. Such schemes have been successfully used in practical photon-counting lifetime detectors. These schemes are attractive because of their efficient utilization of detected photons. However, they suffer from dynamic range problems that arise out of limited counting speeds. Typically, a time-to-voltage converter together with an associated analog (voltage) to digital converter would have a maximum counting rate of around 1 MHz. Also, with this scheme, only one photon can be measured in the interval between laser pulses. These limitations restrict the use of this technique to low light levels when fairly long exposure times are needed in order to obtain sufficient counts for accurate representation of the decay curve. The comparatively large dead-time of this technique can have more insidious consequences. Immediately after the laser pulse, photons will be emitted at the highest rate and therefore more will be preferentially lost at this time because of the dead-time of the detector. This effect will distort the shape of the decay profile.
A spectroscopic microscope system in accordance with the invention is particularly suited to multi-photon spectral imaging to analyze the light emanating from a specimen to identify a fluorophore or provide information on the microenvironment of the fluorophore. It is capable of collecting the desired signal in the presence of background noise with a high degree of accuracy where multiple fluorophores are used and in the presence of endogenous fluorophores. By utilizing spectroscopic analysis of the light emitted from the specimen, the microscope system in accordance with the invention avoids the need for complicated and expensive filters to separate fluorophore signals. The invention is particularly well suited to the imaging and analysis of living specimens by maximum utilization of the light emitted from the specimen and passed to a detector. The spectroscopic analysis is carried out rapidly, allowing real-time analysis of living specimens and avoiding high level of fluorophore exposure to the excitation light.
A spectroscopic microscope system in accordance with the present invention utilizes a light source providing a beam of light that includes a chosen long wavelength that is preferably pulsed in short pulses suitable for multi-photon excitation of a fluorophore. A microscope objective lens receives the beam from the light source and focuses it to a narrow point at which a specimen may be positioned. A dichroic mirror is positioned in a first beam path to direct light from the light source to the objective lens and to direct fluorescent light from a specimen passed back through the objective lens to a second beam path away from the source. A spectral dispersive element is positioned in the second beam path to receive the fluorescent light and to spread the light according to its spectral content. A multi-channel detector having multiple detector elements receives the light spread by the spectral dispersive element and provides output signals indicative of photon events detected by each detector element in the array. A preferred detector element is a photomultiplier tube having multiple detector elements positioned in an array, e.g., 32 elements positioned in side-by-side relation. Discriminator electronics may be coupled to receive the output signals from the multiple channel detector that includes comparators for each channel that compare the output signal of the detector to a threshold and provide an output signal when the threshold is exceeded. A counter is connected in each channel to each comparator to count the pulse outputs from the comparator over a selected period of time. In this manner, the photon events counted in each channel indicate the relative spectral content of the light emitted from the specimen with a high degree of accuracy. The system preferably includes a scanning means to scan the beam in a raster fashion over the objective lens. The discriminator electronics incorporates fast comparators and counters that allow collection of the spectroscopic data on a pixel-by-pixel basis as the excitation light beam is scanned over the specimen.
The excitation of fluorophores in the specimen with relatively long wavelength laser light in short pulses suitable for multi-photon excitation will result in the emission of fluorescent photons at shorter wavelengths than the excitation light. The dichroic mirror may be formed to pass wavelengths of light at selected wavelengths, shorter than that of the beam from the source, which include the wavelengths of the multi-photon absorption fluorescent light emitted from a specimen. Preferably, the light source is a pulsed laser providing light in the range of red to near infrared wavelengths.
In a preferred structure of the microscope system, the second beam path includes a multi-mode optical fiber having an entrance end and an exit end and an objective lens in the second beam path positioned to image the exit aperture of the microscope objective lens onto the entrance end of the optical fiber. The exit end of the optical fiber is positioned to direct a cone of light exiting therefrom onto the dispersive element. A preferred dispersive element is a concave curved holographic diffraction grating, which provides a wide aperture that is especially advantageous for analysis of the relatively low intensity level fluorescent light that can be poorly collimated.
The present invention may also carry out simultaneous fluorescence lifetime imaging utilizing the basic spectroscopic microscope system structure. To provide fluorescence lifetime imaging, the discriminator electronics may include multiple channels, each connected to receive the output of a detector element. Each channel includes a comparator comparing the output of the detector to which it is connected to a reference and providing a pulse output signal when the detector signal is above the threshold. A series of counters is connected to each comparator. The laser light source provides short pulses of light at a selected repetition rate. Means are provided operably connected to the laser source to enable the counters in each series of counters in sequence at selected times after initiation of a light pulse from the laser source. The photon counts determined by each counter in a channel after each light pulse from the laser provide an indication of the photon events as a function of time after the light pulse. The means for enabling the counters preferably includes a tapped delay line connected to receive an enable signal corresponding to the initiation of a light pulse from the laser source. The enable signal is propagated along the delay line, each tap of the delay line being connected to the input of an AND gate the other input of which is connected to the comparator; the output of the AND gate at each tap is preferably connected to one of the counters in the series of counters such that the series of counters are enabled to begin counting in sequence at selected progressively delayed periods of time after initiation of the laser light pulse. Preferably, all of the taps of the tapped delay line are reset to a reset level at a selected time after initiation of the laser light pulse and before the next pulse.
Further objects, features and advantages of the invention will be apparent from the following detailed description when taken in conjunction with the accompanying drawings.